Hemoglobin contrast in magneto-motive optical doppler tomography, optical coherence tomography, and ultrasound imaging methods and apparatus

ABSTRACT

A novel contrast mechanism for imaging blood flow using magneto-motive optical Doppler tomography (MM-ODT), Optical Coherence Tomography, and Ultrasound. MM-ODT, OCT, and ultrasound combined with an externally applied temporally oscillating high-strength magnetic field detects erythrocytes moving according to the field gradient.

FUNDING

This invention was supported by funds from the National Institutes of Health (AR47551, EB002495 and EB002021) and the Texas Advanced Technology Program. The U.S. Government may have certain rights in the invention.

BACKGROUND

The present invention relates in general to the art of medical diagnostic imaging and in particular to imaging blood flow using magneto-motive optical Doppler tomography (MM-ODT), Optical Coherence Tomography, or Ultrasound, which combines an externally applied temporally oscillating high-strength magnetic field with ODT, OCT, or Ultrasound to detect erythrocytes moving according to the field gradient.

The accurate determination of location and flow velocity of moving particles in highly scattering media, such as blood flow, is important for medical diagnostics. While the measurements of blood flow in the coronary arteries is an important aspect in diagnosing coronary artery diseases. Numerous non-invasive approaches have been developed using techniques such as Doppler ultrasound, conventional angiography, laser Doppler flowmetry and magnetic resonance angiography.

One common sensing technique involves the use of ultrasound. Using this technique, ultrasound is directed into the body of the patient and tiny particles such as red blood cells, which are suspended in the blood plasma, scatter the ultrasonic energy back towards the receiver or transducer. The transducer then converts the back-scattered ultrasonic energy into an electrical signal that is processed in some known manner to determine the presence of a flow and an estimate of the flow velocity.

Magnetic resonance imaging (MRI) is based on an imaging technique for magnetically exciting nuclear spins in a subject positioned in a static magnetic field by applying a radio-frequency (RF) signal of the Larmor frequency, and reconstructing an image using MR signals induced by the excitation. MRI is widely applied in clinical medicine because of its capability of clearly depicting the slightest tissue of human brain in vivo.

Magnetic resonance angiography (MRA) provides detailed angiographic images of the body in a non-invasive manner. In conventional MRA, which does not use contrast agents, magnetic resonance signal from flowing blood is optimized, while signal from stationary blood or tissue structures is suppressed. In contrast-enhanced MRA, a contrast agent is injected into the blood stream to achieve contrast between flowing blood and stationary tissue.

The commonly known echo planar imaging (EPI) is a rapid MRI technique, which is used to produce tomographic images at high acquisition rates, typically several images per second. Functional magnetic resonance imaging (fMRI) has been found useful in perfusion and/or diffusion studies and in dynamic-contrast studies, etc. However, images obtained in EPI experiments tend to be vulnerable to an artifact known as “ghosting” or “ghost images.”

Optical coherence tomography (OCT) is a technology that allows for non-invasive, cross-sectional optical imaging of biological media with high spatial resolution and high sensitivity. OCT is an extension of low-coherence or white-light interferometry, in which a low temporal coherence light source is utilized to obtain precise localization of reflections internal to a probed structure along an optic axis. This technique is extended to enable scanning of the probe beam in the direction perpendicular to the optic axis, building up a two-dimensional reflectivity data set, used to create a cross-sectional gray-scale or false-color image of internal tissue backscatter.

OCT uses the short temporal coherence properties of broadband light to extract structural information from heterogeneous samples such as biologic tissue. OCT has been applied to imaging of biological tissue in vitro and in vivo. Systems and methods for substantially increasing the resolution of OCT and for increasing the information content of OCT images through coherent signal processing of the OCT interferogram data have been developed to provide cellular resolution (i.e., in the order of 5 micrometers). During the past decade, numerous advancements in OCT have been reported including real-time imaging speeds.

In diagnostic procedures utilizing OCT, it would also be desirable to monitor the flow of blood and/or other fluids, for example, to detect peripheral blood perfusion, to measure patency in small vessels, and to evaluate tissue necrosis. Another significant application would be in retinal perfusion analysis. Accordingly, it would be advantageous to combine Doppler flow monitoring with the above micron-scale resolution OCT imaging in tissue.

Conventional OCT imaging primarily utilizes a single backscattering feature to display intensity images. Functional OCT techniques process the backscattered light to provide additional information on birefringence, and flow properties. (See for example, Kemp N J, Park J, Zaatar H N, Rylander H G, Milner T E, High-sensitivity determination of birefringence in turbid media with enhanced polarization-sensitive optical coherence tomography, Journal of the Optical Society of America A: Optics Image Science and Vision 2005, 22(3):552-560; Dave D P, Akkin T, Milner T E, Polarization-maintaining fiber-based optical low-coherence reflectometer for characterization and ranging of birefringence, Optics Letters 2003, 28(19):1775-1777; Rylander C G, Dave D P, Akkin T, Milner T E, Diller K R, Welch M, Quantitative phase-contrast imaging of cells with phase-sensitive optical coherence microscopy, Optics Letters 2004, 29(13):1509-1511; de Boer J F, Milner T E, Ducros M G, Srinivas S M, Nelson J S, Polarization-sensitive optical coherence tomography, Handbook of Optical Coherence Tomography, New York: Marcel Dekker, Inc., 2002, pp 237-274.)

Since the ability to characterize fluid flow velocity using OCT was demonstrated by Wang et al., several phase resolved, real-time optical Doppler tomography (ODT) approaches have been reported. (See for example, Chen Z P, Milner T E, Dave D, Nelson J S, Optical Doppler tomographic imaging of fluid flow velocity in highly scattering media, Optics Letters 1997, 22(1):64-66; Wang X J, Milner T E, Nelson J S.

Optical Doppler tomography (ODT) combines Doppler velocimetry with optical coherence tomography (OCT) for noninvasive location and measurement of particle flow velocity in highly scattering media with micrometer-scale spatial resolution. The principle employed in ODT is very similar to that used in radar, sonar and medical ultrasound. ODT uses a low coherence or broadband light source and optical interferometer to obtain high spatial resolution gating with a high speed scanning device such as a conventional rapid scanning optical delay line (RSOD) to perform fast ranging of microstructure and particle motion detection in biological tissues or other turbid media.

To detect the Doppler frequency shift signal induced by the moving particles, several algorithms and hardware schemes have been developed for ODT. The most straightforward method to determine the frequency shift involves the use of a short time fast Fourier transform (STFFT). However, the sensitivity of this method is mainly dependent on the FFT time window, which limits axial scanning speed and spatial resolution when measuring slowly moving blood flow in small vessels that requires high velocity sensitivity. However, a phase-resolved technique can decouple the Doppler sensitivity and spatial resolution while maintaining high axial scanning speed.

In ODT, the Doppler frequency shift is proportional to the cosine of the angle between the probe beam and the scatterer's flow direction. When the two directions are perpendicular, the Doppler shift is zero. Because a priori knowledge of the Doppler angle is not available, and conventional intensity OCT imaging provides a low contrast image of microvasculature structure, detecting small vessels with slow flow rates is difficult. However, the Doppler angle can be estimated by combining Doppler shift and Doppler bandwidth measurements. (See for example, Piao D Q, Zhu Q, Quantifying Doppler Angle and Mapping Flow Velocity by a Combination of Doppler-shift and Doppler-bandwidth Measurements in Optical Doppler Tomography, Applied Optics, 2003, 42(25): 5158-5166, and U.S. Pat. No. 5,991,697 describe a method and apparatus for Optical Doppler Tomographic imaging of a fluid flow in a highly scattering medium comprising the steps of scanning a fluid flow sample with an optical source of at least partially coherent radiation through an interferometer, which is incorporated herein by reference).

The ability to locate precisely the microvasculature is important for diagnostics and treatments requiring characterization of blood flow. Recently, several efforts to increase blood flow contrast mechanisms have been reported including protein microspheres incorporating nanoparticles into their shells, plasmon-resonant gold nanoshells, and use of magnetically susceptible micrometer sized particles with an externally applied magnetic field. (See for example, Lee T M, Oldenburg A L, Sitafalwalla S, Marks D L, Luo W, Toublan F J J, Suslick K S, Boppart S A, Engineered microsphere contrast agents for optical coherence tomography, Optics Letters, 2003, 28(17): 1546-1548; Loo C, Lin A, Hirsch L, Lee M H, Barton J, Halas N, West J, Drezek R. Nanoshell-enabled photonics-based imaging and therapy of cancer. Technology in Cancer Research & Treatment, 2004; 3(1): 33-40; and Oldenburg A L, Gunther J R, Boppart S A, Imaging magnetically labeled cells with magnetomotive optical coherence tomography, Optics Letters, 2005, 30(7): 747-749.)

Wang, et al., “Characterization of Fluid Flow Velocity by Optical Doppler Tomography,” Optics Letters, Vol. 20, No. 11, Jun. 1, 1995, describes an Optical Doppler Tomography system and method which uses optical low coherence reflectrometry in combination with the Doppler effect to measure axial profiles of fluid flow velocity in a sample. A disadvantage of the Wang system is that it does not provide a method to determine direction of flow within the sample and also does not provide a method for generating a two-dimensional color image of the sample indicating the flow velocity and directions within the image.

The use of an externally applied field to move magnetically susceptible particles in tissue has been termed magneto-motive OCT (MM-OCT). Functional magnetic resonance imaging (fMRI) detects deoxyhemoglobin which is a paramagnetic molecule. However, the paramagnetic susceptibility of human tissue is very low compared to other biocompatible agents such as ferumoxides (nanometer sized iron oxide particles). Therefore, it was believed that, other than differentiating relaxation times (T2) between oxygenated and deoxygenated blood, the magnetic field strength required to produce a retarding force on blood flow was well above that of current imaging fields. (See also, for example, Schenck J F., Physical interactions of static magnetic fields with living tissues, Progress in Biophysics and Molecular Biology 2005, 87(2-3):185-204; and Taylor D S, Coryell, C. D., Magnetic susceptibility of iron in hemoglobin. J. Am. Chem. Soc. 1938, 60:1177-1181.)

The present invention is a novel extension of MM-OCT to image hemoglobin in blood erythrocytes. In one aspect of the invention, the approach requires no exogenous contrast agent to detect blood flow and location.

SUMMARY OF THE INVENTION

The invention is a method and apparatus of imaging a blood flow using Doppler optical coherence tomography, which comprises an externally applied temporally oscillating high-strength magnetic field with a Doppler optical tomography system to detect hemoglobin moving according to the magnetic field gradient.

In another embodiment, the invention comprises a solenoid cone-shaped ferrite core with an extensively increased magnetic field strength at the tip of the core. In a further embodiment, the invention comprises focusing the magnetic force on targeted samples.

Another embodiment is an apparatus for imaging blood flow, comprising: (1) a magnetomotive optical Doppler tomography (MM-ODT) imaging system; (2) an oscillating magnetic field applied to the moving blood flow; (3) a light source (4) a rapid scanning optical delay line in a reference arm and aligned such that no phase modulation is generated when the group phase delay is scanned at 4 kHz; and (5) a hardware in-phase and a quadrature demodulator with high bandpass filters.

Another aspect of the invention is a method for imaging a blood flow, comprising applying a magnetic field to the blood flow, wherein the blood flow comprising a plurality of hemoglobin molecules and wherein the magnetic field interacts with the hemoglobin molecules to cause a change in the blood flow; and detecting the blood flow by detecting the change in the blood flow caused by the interaction with the hemoglobin molecules, wherein the change is detected using a magnetomotive optical Doppler tomography imaging system.

Another embodiment of the method for imaging a blood flow comprises temporally oscillating the magnetic field.

Another embodiment of the invention is an apparatus for imaging a blood flow comprising a magnetic field generator for applying a magnetic field to the blood flow, wherein the blood flow comprises hemoglobin molecules; and an ultrasound detection system for detecting the blood flow while it is in the presence of the magnetic field.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic diagram of the MM-ODT system.

FIG. 2 is a schematic diagram of the probe beam, flow sample and solenoid coil.

FIGS. 3 a and 3 b are OCT and ODT images of a stationary turbid solution without an external magnetic field, respectively. FIG. 3 c and FIG. 3 d are OCT and ODT images with a 50 Hz magnetic field, respectively. The white bar represents 200 μm, accordingly.

FIGS. 4 a-4 d are M-mode ODT images of the diluted deoxygenated blood flow (18% hematocrit) without and with an external magnetic field. FIG. 4 a and FIG. 4 b are ODT images of 5 mm/s blood flow without and with a 5 Hz magnetic field, respectively. FIG. 4 c and FIG. 4 d are ODT images of 30 mm/s blood flow without and with a 50 Hz magnetic field, respectively. The Black bar indicates 200 μm, accordingly.

FIG. 5 a and FIG. 5 b are the Doppler frequency shift profiles without an external magnetic field, and with a 50 Hz magnetic field, respectively.

FIG. 6 is a block diagram illustrating an exemplary phase sensitive OCT system.

FIG. 7 is a graph of the optical path length change in the DP-OCT system.

FIG. 8 is a schematic diagram of Magneto Motive Ultrasound for Blood.

FIG. 9 is an ultrasound image of blood exposed to an oscillating magnetic field.

FIG. 10 is a schematic diagram of Ultrasonography and OCT coupled with magnetic field generator.

DETAILED DESCRIPTION OF THE INVENTION

Magneto-Motive Optical Doppler Tomography (MM-ODT) for improved Doppler imaging of blood flow using an external oscillating magnetic field is described below. By introducing mechanical movement of red blood cells (RBC's) during blood flow by a temporally oscillating high-strength magnetic field, MM-ODT allows imaging of blood flow and velocity. The controlled and increased Doppler frequency in MM-ODT provides an investigational tool to study in vivo blood transport, as shown in the article Hemoglobin Contrast in Magnetomotive Optical Doppler Tomography, Opt. Lett. 31, 778-780 (2006), herein incorporated by reference.

The microstructure of the blood flow and flow velocity information are all encoded in the interferogram of a Doppler OCT system. It should be readily apparent to those skilled in the optical arts, that different OCT systems and different OCT information can be used to determine the Doppler frequency shift. It is not intended to suggest any limitation as to the scope or functionality with different OCT architectures or optical information used with an oscillating magnetic field, such as time domain Doppler OCT and spectral domain Doppler OCT. Time domain OCT requires a mirror in the reference arm scanning at a constant velocity, while spectral domain OCT includes swept source OCT and Fourier domain OCT. An example of time-domain Doppler OCT is provided below.

A schematic of the MM-ODT apparatus 10 is shown in FIG. 1. The OCT light source 12 comprises a super luminescent diode, which is used as the low coherence light source. In one embodiment, light source 12 is centered at 1.3 μm with a bandwidth of 90 nm. Light from source 12 is coupled into a single-mode optical fiber based interferometer 20 by circulator 14, where the interferometer 20 can provide 1 mW of optical power at the sample 60. Light is split into a reference arm 22 and sample arm 24 by a 2×2 splitter 16. A rapid-scanning optical delay (RSOD) line 26 is coupled to the reference arm 22. In one embodiment, the rapid-scanning optical delay line 26 is aligned such that no phase modulation is generated when the group phase delay is scanned at ˜4 kHz. In the sample arm 24, a collimated beam 36 is redirected to sample 60 by two galvanometers 30 that permit three-dimensional scanning. In one embodiment, the galvanometers can be an X-scanner and a Y-scanner. The sample 60 can be internal or external to the body, where the probe beam is focused by an objective lens 32. In one embodiment, the objective lens 32 yields a 10-μm spot at the focal point. Phase modulation can be generated using an electro-optical waveguide phase modulator, which can produce a carrier frequency (˜1 MHz). And the magnetic field generator 100 is in proximity to the sample 60.

A dual-balanced photodetector 34 is coupled to the 2×2 splitter 16 and the circulator 14. The photodetector 34 of a 80 MHz bandwidth reduces the light source noise from the OCT interference signal. A hardware in-phase and quadrature demodulator 40 with high/bandpass filters 42 and low/bandpass filters 44 improves imaging speed. Doppler information was calculated with the Kasai autocorrelation velocity estimator. Labview software 50 (National Instruments, Austin, Tex.) is coupled to the MM-ODT system with a dual processor based multitasking scheme. The maximum frame rate of the MM-ODT system 10 was 16 frames per second for a 400×512 pixel sized image. The Doppler frequency shift can be determined with the use of a short time fast Fourier transform (STFFT). Alternatively, a phase-resolved technique can determine the Doppler frequency shift to decouple the Doppler sensitivity and spatial resolution while maintaining high axial scanning speed. Alternatively, differential phase optical coherence tomography (OCT) or spectral domain phase-sensitive OCT can be used to determine the Doppler frequency shift, as readily determined by one skilled in the optical arts.

FIG. 2 shows an example of the magnetic field generator 80 with a capillary tube 90. The magnetic field generator 80 includes a solenoid coil 82 (Ledex 4EF) with a cone-shaped ferrite core 84 at the center and driven by a current amplifier 86 supplying up to 960 W of power. The magnetic field generator 80 can be placed underneath the sample 60 during MM-ODT imaging. The combination of the ferrite core 84 and solenoid coil 82 using a high power operation dramatically increases the magnetic field strength (B_(max)=0.14 Tesla) at the tip of the core 82 and also focuses the magnetic force on the targeted samples 60. The sinusoidal current can vary the magnetic force applied to the capillary tube 90 in order to introduce movement of magnetic fluids, which include red blood cells that contain hemoglobin. In one embodiment, the probe beam is oriented parallel to the gradient of the magnetic field's strength.

The material parameter characterizing magnetic materials, including biological tissue, is the magnetic volume susceptibility, χ. Magnetic volume susceptibility is dimensionless in SI units and is defined by the equation M=χH, where M is the magnetization at the point in question and H is the local density of the magnetic field strength. Hemoglobin's high iron content, due to four Fe atoms in each hemoglobin molecule, and the large concentration of hemoglobin in human red blood cells give Hemoglobin magneto-motive effects in biological tissue. The magnetic volume susceptibility of the hemoglobin molecule consists of a paramagnetic component due to the electron spins of the four iron atoms. The paramagnetic susceptibility is given by the Curie Law,

$\begin{matrix} {\chi = \frac{\mu_{o}{N_{p}\left( {\mu_{eff}^{2}\mu_{B}^{2}} \right)}}{3{kT}}} & (1) \end{matrix}$ where μ_(o) is permeability of free space and has the value 4π×10⁷ H/m, N_(p) is the volume density of paramagnetic iron atoms in hemoglobin, N_(p)=4.97×10²⁵ iron atoms/m³, μ_(eff) is the effective number of Bohr magnetons per atom reported as 5.35, and the Bohr magneton, μ_(B)=9.274×10⁻²⁴ J/T, and Boltzmann's constant, k=1.38×10⁻²³ J/K, and T is the absolute temperature (K). The calculated susceptibility of a RBC is about 11×10⁻⁶ assuming a 90% concentration of hemoglobin per RBC. The calculated susceptibility of a RBC is dependent on the oxygenation of the hemoglobin. The calculations can be adjusted accordingly, depending on the oxygenation of the RBC, which can be measured by known techniques.

A RBC placed in a magnetic field gradient experiences forces and torques that tend to position and align it with respect to the field's direction. The magnetic force, in the direction of the probing light z, is given by

$\begin{matrix} {{F_{z} = {{m_{RBC}\frac{\delta^{2}{z(t)}}{{dt}^{2}}} = {\frac{\delta\; U}{\delta\; z} = {\frac{\chi\; V}{\mu_{o}}B\frac{\delta\; B}{\delta\; z}}}}},} & (2) \end{matrix}$ where V is the particle volume, B is the magnitude of the magnetic flux density, and Δχ is the difference between the susceptibility of the particle and the surrounding medium. The displacement [z(t)] of an RBC driven by a time varying magnetic flux density can be included in the analytic OCT fringe expression, I_(f),

$\begin{matrix} {{I_{f}a\; 2\sqrt{I_{r}I_{s}{\exp\left\lbrack {{\mathbb{i}}\left( {{2\pi\; f_{o}t} + {\frac{4\pi\;{z(t)}}{\lambda_{o}}{z(t)}}} \right)} \right\rbrack}}},} & (3) \end{matrix}$ where I_(R) and I_(S) are the back scattered intensities from the reference and sample arms, respectively, f_(o) is the fringe carrier frequency, λ_(o) is the center wavelength of the light source, and z(t) is the RBC displacement. Integration of all forces (magnetic, elastic, and viscous) on the RBC gives displacement, z(t)=A cos(4πf_(m)t) where A is a constant in units of length and f_(m) is the modulation frequency of the magnetic flux density. In free space, the displacement, z(t), is dominated a constant acceleration which can be, however, ignored in confined models (i.e. blood vessel or capillary tube) with assumptions that, first, the probing area is much smaller than magnetic field area, and that secondly probing time starts after steady states of inner pressures. Expansion of the right-hand side of Eq. (3) using Bessel functions gives

$\begin{matrix} {I_{f}\alpha\; 2\sqrt{I_{R}I_{S}}\left( {{\sum\limits_{k = 0}^{\infty}\left( {{J_{k}(m)}{\exp\left( {{\mathbb{i}}\; k\; 4\pi\; f_{m}t} \right)}} \right)} + {\sum\limits_{k = 0}^{\infty}\left( {\left( {- 1} \right)^{k}{J_{k}(m)}{\exp\left( {{- {\mathbb{i}}}\; k\; 4{pf}_{m}t} \right)}} \right)}} \right){\exp\left( {{\mathbb{i}2\pi}\; f_{o}t} \right.}} & (4) \end{matrix}$ where J_(k)(m) is the Bessel function of the first kind of order k for argument m which is 4πA/λ_(o). The amplitude of the k^(th) sideband is proportional to J_(k)(m). In coherent detection, the fraction of optical power transferred into each of the first order sidebands is (J_(l)(m))², and the fraction of optical power that remains in the carrier is (J_(o)(m))².

EXAMPLE 1

MM-ODT Imaging of the Doppler Shift of Hemoglobin by Applying an Oscillating Magnetic Field to a Moving Blood Sample

M-mode OCT/ODT images of a capillary glass tube filled with a stationary turbid solution with and without an external magnetic field as a control sample were recorded, as shown in FIGS. 3-4. A 750 μm-inner diameter glass capillary tube 200 was placed perpendicularly to the probing beam 70, as shown in FIG. 2. Fluids used for flow studies were injected through the tube at a constant flow rate controlled by a dual-syringe pump (Harvard Apparatus 11 Plus, Holliston, Mass.) with ±0.5% flow rate accuracy. The turbid solution was a mixture of deionized water and 0.5-gm latex microspheres (μ_(s)=5 mm⁻¹). The magnetic flux density and its frequency were approximately 0.14 T and 50 Hz, respectively. M-mode OCT/ODT images were acquired for 100 ms per frame. FIGS. 3 a and 3 b show M-mode OCT and ODT images without any external magnetic field, respectively. The ODT image in FIG. 3 b contains small random phase fluctuations due to ambient vibration through the optical path. FIGS. 3 c and 3 d show M-mode OCT and ODT images with a 50 Hz external magnetic field, respectively. No distinguishable Doppler shift could be observed in the ODT image FIG. 3 d indicating no interaction between the external magnetic field and the moving microspheres.

Deoxygenated blood was extracted from the vein of a human male's left arm, and diluted with saline. During preparation, blood was not exposed directly to air so as to remain deoxygenated. To simulate flow, blood was injected through the capillary tube 200 by a syringe pump at a relatively constant flow rate. As FIG. 4 shows, the oscillating Doppler frequency shift, resulting from RBC movement, could be observed at two different flow rates (5 and 30 mm/s). Because the flow direction was nearly perpendicular to the probing beam no significant Doppler frequency shift was distinguishable at the 5 mm/s flow rate FIG. 4 a without any external magnetic field. In the case of the high blood flow rate of 30 mm/s, as shown at FIG. 4 c, the Doppler frequency shift caused by the blood flow could be observed. And for maximum contrast enhancement, the probe beam can be directed parallel to the gradient of the magnetic field's strength. However, application of a 50 Hz magnetic field increased the Doppler contrast of blood at both the slow and fast flow rates as shown at FIGS. 4 b and 4 d. The high flow rate of 30 mm/s gives a higher contrast image than the low flow rate image, but the Doppler frequency shift of the former as a function of depth is less homogeneous than the latter, which is indicative of perturbation by blood flow. The same blood was diluted to 5% hematocrit (HCT), but no RBC movement could be observed below 8% HCT.

Doppler frequency shift profiles were calculated from the ODT images by averaging 20 lines at a selected depth indicated by horizontal arrows, as shown in FIGS. 4 a and 4 b. FIG. 5 a indicates no significant Doppler frequency shift over a 100 ms time period, whereas FIG. 5 b displays ±200 to 300 Hz Doppler frequency shifts oscillating 20 times over 100 ms (200 Hz).

The Doppler frequency shift indicates that RBC's physically move into and away from the incident light while passing through the external magnetic field depending on whether their magnetic properties are paramagnetic or diamagnetic, as shown in FIG. 4, and that the 200 Hz oscillation of the Doppler frequency shift correlates with the 50 Hz magnetic field, as shown in FIG. 5. Frequencies 4 times higher than that of the external magnetic field (f_(m)) can be observed. According to Eq. 2, the frequency of the force on paramagnetic targets was twice that of the magnetic flux density; therefore, a 50 Hz B field displaces the targets at 100 Hz. Although the fringe signal (Eq. 4) contains harmonics at frequency (2f_(m)), the modulation frequency, f_(m), was set so that the second-order sideband (4f_(m)) was dominant as shown in FIG. 5 b. The particle motion can not be described as a pure sinusoidal function even if the modulated magnetic field is sinusoidal, due to the numerous forces that contribute to the motion within the field such as gravity, concentration gradient, and colloidal dispersion.

The invention is a new investigational tool to study in vivo blood transport and the first implementation of MM-ODT for improved Doppler imaging of blood flow using an external oscillating magnetic field introducing a mechanical movement of RBC's during blood flow by a temporally oscillating high-strength magnetic field. MM-ODT to allow imaging of tissue function in a manner similar to functional magnetic resonance images (f-MRI) of deoxygenated blood in organs, when the sample arm of the MM-ODT system is coupled to a probe (not shown). Such probes are generally known in the arts, such as endoscopic probes, catheter probes, and the like.

Alternatively, the MM-ODT can be used for Port-Wine vessel mapping and Skin Cancer vessel mapping. The MM-ODT can be used to detect blood vessel location and size for cancer and port-wine stains, since these conditions are characterized by blood vessel growth and increases in hemoglobin content. Accordingly, other blood vessel detection for tissue abnormality identification can be envisioned with this invention. Generally, the MM-ODT can be used wherever blood flow detection is necessary in operations, chemotherapy, hemodialysis, and the like.

A spectral domain phase sensitive OCT system 100 can be used to image the blood flow and determine velocity information, when an oscillating magnetic field is applied to the blood flow, as shown in FIG. 6. Spectral domain phase sensitive OCT generally uses a broadband light source in a general interferometric setup, where the mirror in the reference typically does not move or does not require a rapid scanning delay line. A spectral interferogram is detected, in which each A-scan is entirely encoded. A Fourier transformation of the A-scan and velocity distribution of the blood flow can be extracted by known methods. In spectral domain Doppler OCT, the blood flow profile is the envelope of a calculated A-scan through Fourier transformation on the spectral interferogram. Velocity information is encoded in the phase of the interferogram. The Doppler velocity can be extracted by measuring the phase shift between two successive calculated A-scans in spectral domain Doppler OCT.

This OCT system 100 is only an example of one OCT imaging modality which can be used to image blood flow with a temporally oscillating magnetic field, and is not intended to suggest any limitation on the scope of OCT architectures applicable to the invention. Generally, the OCT system 100 includes a general-purpose computing device in the form of a computer 101 and includes a magnet control 114 and a magnet 116.

Light energy is generated by a light source 117. The light source 117 can be a broadband laser light source coupled into optical fiber emitting light energy over a broad range of optical frequencies. The wavelength range can be from about 400 nanometers to about 1400 nanometers. Longer wavelengths (>600 nm) can be used for deeper scanning. Preferably, the light source emits light having a wavelength near the infrared spectrum to identify hemoglobin for OCT imaging, which places hemoglobin in motion and increases optical scattering of the hemoglobin. The light energy can be emitted over a multiplicity of optical wavelengths, frequencies, and pulse durations to achieve OCT imaging. As used herein, optical fiber can refer to glass or plastic wire or fiber. Optical fiber is indicated on FIG. 6 as lines connecting the various blocks of the figures. Where light energy is described as “passing,” “traveling,” “returning,” “directed,” or similar movement, such movement can be via optical fiber.

A fraction of the generated light energy passes from the light source 117 into an optical spectrum analyzer 118. The optical spectrum analyzer 118 measures optical frequency as the light energy is emitted from the light source 117 as a function of time. The optical spectrum analyzer 118 samples a portion of the light emitted by the light source 117. The optical spectrum analyzer 118 monitors the power spectral density of light entering the splitter 119. The remaining fraction of light energy from the light source 117 passes into a splitter 119. The splitter 119 can be a device with four ports, with Port 1 allowing light energy to enter the splitter 119. Ports 2 and 3 allow light energy to leave and re-enter the splitter 119 to the reference reflector 120 and OCT probe 122, respectively. Port 4 allows light energy to leave the splitter 119 to coupling lens 124. The splitter 119 couples the light into Port 1. The splitter 119 divides the light according to a pre-determined split ratio selected by a user. For example, the split ratio can be 50/50 wherein half of the light energy entering the splitter 119 at Port 1 exits the splitter 119 through Port 2 and half exits the splitter 119 through Port 3. In another example, the split ratio can be 60/40 wherein 60% of the light energy passes through Port 2 and 40% of the light energy passes through Port 3.

A fraction of the light energy (determined by the split ratio) that exits the splitter 119 through Port 2 travels to a reference reflector surface 120. The light energy is reflected from the reference reflector surface 120 back to the splitter 119 into Port 2. The reference reflector 120 can be a planar metallic mirror or a multilayer dielectric reflector with a specified spectral amplitude/phase reflectivity. The remaining fraction of light that entered splitter 119 through Port 1 exits splitter 119 through Port 3 and enters an OCT probe 122. The OCT probe 122 can be a turbine-type catheter as described in Patent Cooperation Treaty application PCT/US04/12773 filed Apr. 23, 2004 which claims priority to U.S. provisional application 60/466,215 filed Apr. 28, 2003, each herein incorporated by reference for the methods, apparatuses and systems taught therein. The OCT probe 122 can be located within a subject 121 to allow light reflection off of subject's 121 blood flow.

The light energy that entered OCT probe 122 is reflected off of the blood flow of subject 121 once an oscillating magnetic field has been temporally applied by magnet 116. The reflected light energy passes back through the OCT probe 122 into the splitter 119 via Port 3. The reflected light energy that is returned into Port 2 and Port 3 of the splitter 119 recombines and interferes according to a split ratio. The light recombines either constructively or destructively, depending on the difference of pathlengths. A series of constructive and destructive combinations of reflected light create an interferogram (a plot of detector response as a function of optical path length difference). Each reflecting layer from the subject 121 and the blood flow will generate an interferogram. The splitter 119 can recombine light energy that is returned through Port 2 and Port 3 so that the light energies interfere. The light energy is recombined in the reverse of the split ratio. For example, if a 60/40 split ratio, only 40% of the light energy returned through Port 2 and 60% of the light energy returned through Port 3 would be recombined. The recombined reflected light energy is directed out Port 4 of the splitter 119 into a coupling lens 137. The coupling lens 137 receives light from the output of the splitter 119 and sets the beam etendue (beam diameter and divergence) to match that of the optical spectrometer 138. The coupling lens 137 couples the light into an optical spectrometer 138. The optical spectrometer 138 can divide the recombined reflected light energy light into different optical frequencies and direct them to different points in space which are detected by a line scan camera 139. The line scan camera 139 performs light to electrical transduction resulting in digital light signal data 108. The digital light signal data 108 is transferred into the computer 101 via the OCT input interface 111. Interface between the line scan camera 139 and computer 101 can be a Universal Serial Bus (USB), or the like. The digital light signal data 108 can be stored in the mass storage device 104 or system memory 112 and utilized by the image construction software 106 and the Labview image construction software 107.

The image construction software 106 can generate an image of the blood flow of subject 121 from the light signal data 108, by receiving light signal data 108 generating amplitude and phase data. The amplitude and phase data (optical path length difference (cτ) or optical time-delay (τ)) can be separated into discrete channels and a plot of intensity vs. depth (or amplitude vs. depth) can be generated for each channel. Such plot is known as an A-scan, where the composition of all the A-scans can comprise one image. And movement image construction software 107 generates an image of the movement of the hemoglobin from the light signal data 108. The movement image construction software 107 receives light signal data 108 for at least two successive sweeps of the light source 117 or the light source performs a Fourier transform on the light signal data 108 generating amplitude and phase data.

Optionally, additional information can be extracted from the light signal data to generate additional images. The light signal data can be further processed to generate a Stokes parameter polarimetric image when used in conjunction with polarization detectors and polarizing lenses to extract polarization data from the light signal 108, as readily known to one skilled in the art of optical coherence tomography. The differential phase OCT image data is shown in FIG. 7, indicating the optical path difference with the magnetic field. Filters can be added to reduce the noise of the signal generated by the magnetic field.

Alternatively, the phase sensitive OCT system 100 can be configured for swept source OCT, which is a different type of spectral domain OCT. In swept source OCT, a tunable laser source replaces the broadband laser light source 117. The scanning rate can be at wavelengths of 800 nm-1500 nm. Also, the reference reflector surface 120 is in-line with sample path 120. The optical spectrometer 125 and line scan camera 126 are replaced with a general photodetector.

In another embodiment, an enhanced detection of cancer with ultrasound imaging 200 is provided. Ultrasonography is the ultrasound-based diagnostic imaging technique used to visualize muscles and internal organs, their size, structures and any pathological lesions. “Ultrasound” applies to all acoustic energy with a frequency above human hearing (20,000 Hertz or 20 kilohertz). Typical diagnostic sonography scanners operate in the frequency range of 2 to 13 megahertz, hundreds of times greater than this limit. The choice of frequency is a trade-off between the image spatial resolution and penetration depth into the patient, with lower frequencies giving less resolution and greater imaging depth. Doppler ultrasonography uses the Doppler Effect to assess whether blood is moving towards or away from a probe, and its relative velocity. By calculating the frequency shift (υ_(D)) of a particular sample volume, for example a jet of blood flow over a heart valve, its speed and direction can be determined and visualized. Ultrasonagraphy and Doppler Ultrasonagraphy can best be understood by S. A. Kana Introduction to physics in modern medicine, Taylor & Francis, (2003). The basic physics of the Doppler effect involving acoustic and electromagnetic waves of OCT is similar and many of the signal processing techniques (hardware and software) used to estimate the Doppler shift is analogous.

In one embodiment of the invention, an ultrasound probe 212 is coupled with the magnetic field generator 100, as shown in FIG. 8. Ultrasound 210 is directed into the body of the patient by known techniques and the moving red blood cells backscatter the ultrasonic energy back towards the transducer of the ultrasound. The oscillating magnetic field generated by the magnetic field generator 100 increases the contrast of the ultrasonic energy 210 received from the red blood cells. The transducer then converts the back-scattered ultrasonic energy 210 into an electrical signal that is processed in some known manner to determine an estimate of the flow. An enhanced ultrasound image is produced, as displayed in FIG. 9.

In one example, a rectal ultrasound probe is coupled with a magnet to evaluate the prostate gland for cancer. Currently, ultrasound is used for prostate cancer screening; however, the approach provides poor sensitivity and specificity. Yet, all cancers are known in the art to be highly vascular, so then the application of the magnetic field by generator 100 coupled with ultrasound enhances the contrast available from the endogenous RBC's in the prostate for cancer detection at an earlier stage. It is generally known in the art that cancers have enhanced metabolic properties compared to normal tissues, so then cancerous cells have higher oxygen content from hemoglobin and a greater concentration of deoxygenated hemoglobin compared to normal tissues. An exemplary ultrasound image for prostate cancer screening is shown in FIG. 9.

MM-ODT, OCT, and ultrasound coupled with an oscillating magnetic field all can be used in clinical management of patients who need microvasculature monitoring, as shown in FIG. 10. The images could monitor and determine tissue profusion and viability before, during and after operation procedures. For example, the images could be used to detect oxygenated and deoxygenated blood supply, detect blood vessel location, sizes of cancer, vascular tissue abnormality identification, and port-wine stain. Alternatively, the images could monitor photodynamic therapy or evaluate cranial injuries. While medical applications of the invention have been described, the embodiments of the invention are applicable to any circumstance where image contrast is needed for fluids comprising endogenous metallic compositions. MM-ODT, OCT, and ultrasound can be used in various combinations to detect vascular occurrences.

While the invention has been described in connection with various embodiments, it will be understood that the invention is capable of further modifications. This application is intended to cover any variations, uses or adaptations of the invention following, in general, the principles of the invention, and including such departures from the present disclosure as, within the known and customary practice within the art to which the invention pertains. 

1. A system for imaging blood vessels, comprising: a magnetic field generator configured to produce an oscillating magnetic field to be applied to the blood vessels to be imaged, wherein the oscillating magnetic field being greater than or equal to 0.14 T at a time during each oscillatory cycle and the magnetic field is oscillating at a rate greater than or equal to 50 Hz to magnetically move hemoglobin; and an optical coherence tomography apparatus configured to detect magnetic movement of hemoglobin within the blood vessels.
 2. The system of claim 1, further comprising an ultrasound imaging system.
 3. The system of claim 1, wherein the magnetic field generator comprises a solenoid coil with a ferrite core.
 4. The system of claim 3, wherein the optical coherence tomography apparatus further comprises a sample arm, and wherein a probe is coupled to the sample arm of the optical coherence tomography apparatus and the magnet.
 5. The system of claim 4, wherein the optical coherence tomography apparatus is a magneto-motive optical Doppler tomography imaging system for detecting the blood flow affected by the magnetic field of an at least 100 Hz Doppler frequency shift.
 6. The system of claim 5, wherein the magneto-motive optical Doppler tomography system comprises: a light source generating light energy; an interferometer coupled to the light energy, wherein the interferometer includes a reference and sample light paths coupled to a light splitter; a modulator coupled to the interferometer for modulating the optical path length difference in the reference arm and the sample arm; a scanner coupled to the sample arm for scanning a biological sample; a rapid scanning optical delay line coupled to the reference arm; a photodetector coupled to the interferometer for detecting backscattered radiation received by the interferometer from the scanner to detect interference fringes; and a processor for processing the reflected light energy from the reference arm and a magnetic resonance signal reflected off of the moving blood flow to produce a tomographic image and a tomographic flow velocity image.
 7. The system of claim 6, wherein the magneto-motive optical Doppler tomography system includes the scanner with at least two galvanometers to permit three dimensional scans.
 8. The system of claim 7, wherein the magneto-motive optical Doppler tomography system includes a dual-balanced photodetector.
 9. The system of claim 8, wherein the magneto-motive optical Doppler tomography system includes a circulator coupled to the interferometer and the photodetector.
 10. The apparatus of claim 4, wherein the optical coherence tomography system for detecting blood flow is a spectral domain phase sensitive optical coherence tomography system.
 11. The apparatus of claim 10, wherein the phase sensitive optical coherence tomographic imaging modality is a swept source phase sensitive optical coherence tomography system.
 12. The apparatus of claim 2, wherein the optical coherence tomography apparatus detects the displacement of red blood cells driven by the oscillating magnetic field.
 13. The apparatus of claim 4, wherein the probe is oriented parallel to the gradient of the magnetic field.
 14. A system for imaging hemoglobin within the blood vessels, comprising: a magnetic field generator configured to produce an oscillating magnetic field to be applied to the blood vessels to be imaged, the oscillating magnetic field having a strength and oscillation sufficient to induce a detectable magnetic movement of hemoglobin having a magnetic susceptibility of about 11×10⁻⁶; and an optical coherence tomography apparatus configured to detect magnetic movement of hemoglobin within the blood vessels.
 15. The system of claim 14, further comprising an ultrasound imaging system.
 16. The system of claim 14, wherein the magnetic field generator comprises a solenoid coil with a ferrite core.
 17. The system of claim 14, wherein the optical imaging apparatus further comprises a sample arm, and wherein a probe is coupled to the sample arm of an optical coherence tomography apparatus and the magnetic field generator.
 18. The system of claim 14, wherein the optical imaging apparatus is a magneto-motive optical Doppler tomography imaging system for detecting the blood flow affected by the magnetic field.
 19. The system of claim 18, wherein the magneto-motive optical Doppler tomography system further comprises: a light source generating light energy; an interferometer coupled to the light energy, wherein the interferometer includes a reference and sample light paths coupled to a light splitter; a modulator coupled to the interferometer for modulating the optical path length difference in the reference arm and the sample arm; a scanner coupled to the sample arm for scanning a biological sample; a rapid scanning optical delay line coupled to the reference arm; a photodetector coupled to the interferometer for detecting backscattered radiation received by the interferometer from the scanner to detect interference fringes; and a processor for processing the reflected light energy from the reference arm and a magnetic resonance signal reflected off of the moving blood flow to produce a tomographic image and a tomographic flow velocity image.
 20. The system of claim 19, wherein the magneto-motive optical Doppler tomography system includes the scanner with at least two galvanometers to permit three dimensional scans.
 21. The system of claim 19, wherein the magneto-motive optical Doppler tomography system includes a dual-balanced photodetector.
 22. The system of claim 19, wherein the magneto-motive optical Doppler tomography system includes a circulator coupled to the interferometer and the photodetector.
 23. A system for imaging hemoglobin comprising: a magnetic field generator configured to produce an oscillating magnetic field to hemoglobin, wherein the oscillating magnetic field being greater than or equal to 0.14 T at a time during each oscillatory cycle and the magnetic field is oscillating at a rate greater than or equal to 50 Hz to magnetically move hemoglobin; and an imaging apparatus configured to detect magnetic movement of hemoglobin.
 24. The system of claim 23, wherein the imaging system is an optical coherence tomography apparatus.
 25. The system of claim 23, further comprising an ultrasound imaging system. 